It is known in the field of nuclear magnetic resonance (NMR) imaging that it is desirable to reduce image scan times. In order to generate an NMR signal the net magnetic moment (M) of a sample comprising many nuclei is manipulated with suitable RF pulses. A variety of commonly used pulse sequences are known such as free induction decay (FID), inversion recovery and the well-known spin-echo sequence. In traditional magnetic resonance imaging systems a uniform magnetic field B0 is applied across the physical entity being imaged. By physical entity it is meant any object or other entity that is the subject of imaging such as for example a part of the human or animal body, fluid flow in the human or animal body or an industrial fluid flow. Typically a super conducting magnet is employed and uniformity in the magnetic field is achieved using a plurality of shim coils, which superimpose small high-order magnetic fields on B0, proportional to x, y, z, z2, z2y, etc.
The main components of a prior art magnetic resonance imaging system, as used in the medical field, are illustrated schematically in FIG. 1. The x, y, z gradient coils 101 and power supply 102 are provided to apply a gradient field along the length of an imaging chamber positioned in the field of the super conducting magnet 103, the magnet being powered by a dedicated power supply means 104. The gradient field thereby changes the field along the length (z axis) of the imaging chamber. A radio-frequency (RF) coil system 105 is provided along with transmitter circuitry and a power supply 106 to generate a second, lower magnitude, magnetic field of flux density B1 oriented in the x y plane and rotating at the Larmor angular frequency (_∘) The net magnetic moment (M) of a region, along the length of the imaging chamber, corresponding to the resonance frequency at the given region will thus experience a second torque. The angle that M moves through from the z-axis will depend on the magnitude and duration of B1 The flux density B1 has a frequency in the radio frequency range for NMR measurements and a 90° RF pulse is one that turns M through 90° into the x y plane. The x y component of M is a rotating RF field that can be detected with a suitable coil (e.g. coil 104 or a further coil) and appropriately configured electronic signal detection circuitry 107. Additionally in the known MRI apparatus there are provided shim coils 108 powered by power supply means 109.
The known imaging apparatus comprising components 101 to 109 is controlled by a central processing unit 110 (CPU). However a special processor, the sequence controller 111, is configured to effect given imaging sequences to be undertaken in respect of a particular physical entity being imaged. Sequence controller 111 thus affects control over the gradient coils 101, shim coils 108 and the RF coil (or system of coils) 105. The RF coil system may suitably comprise a multiple channel B1 coil array with post-processing software for reconstructing from reduced k-space data. Upon an image of a physical entity being captured the central processing unit 110 is configured to process the received image data in accordance with the predetermined-processing regime implemented. Typically a Fourier transformation is performed upon the data and the resultant image displayed on a display such as display 112. Alternatives to Fourier transformation processing exist such as, for example, the method of ‘projection reconstruction’ which has become used less and less since the first commercial scanners were in use.
FIG. 2 schematically illustrates a known magnetic resonance imaging chamber 201, comprising sample placement region 202, super conducting coil arrangement 203 and super conducting magnet power supply connection means 204. As shown in FIG. 2 a uniform magnetic field B0 exists along the longitudinal axis (z axis) of the super conducting magnet coil arrangement 203.
To obtain an MR image of a slice through a given physical entity, the physical entity must be positioned in the uniform magnetic field to enable resonance to occur. Following placement in the magnetic field a radio frequency pulse is applied at the frequency that corresponds to the given location for which an image slice is required. The physical entity, upon being exposed to the radio frequency pulse, thereby absorbs and re-emits the signal, the signal being detected by the electronic detection circuitry 107 shown schematically in FIG. 1. As discussed above, the frequency at which resonance occurs depends on the magnetic field at a particular location. Thus if the magnetic field changes linearly along its length then the resonant frequency for the protons of a given sample also changes linearly. Thus to excite a particular slice in a given sample physical entity a RF pulse must be generated at the correct offset from the resonant frequency (F) for the protons (0, F, 2F, 3F, 4F etc.) which corresponds to the particular slice of the physical entity for which an image is required. FIG. 3 schematically illustrates a uniform magnetic field generated by the prior art-imaging chamber shown in FIG. 1 and FIG. 2 and additionally depicts an applied gradient 302; the position of a slice in the object being imaged is indicated at 303 as being equivalent to an RF pulse of resonance frequency 3F.
In order to detect a nuclear magnetic resonance signal it is essential that several conditions are met as follows:    1. A sample containing nuclei is polarised in a static magnetic field known as the B0 field creating a net nuclear magnetization.    2. A radio frequency field pulse is applied with a certain frequency, amplitude and duration orthogonally to B0 in order to perturb the nuclear magnetization away from alignment with the main magnetic field. This field is conventionally known as the B1 field and is an oscillating field at a specific frequency determined by the property of the specific nuclear spin species under investigation, known as the magnetogyric ratio. The specific frequency required is related to the static magnetic field strength through a linear relationship known as the Larmor equation, the proportionality constant between frequency and field being the magnetogyric ratio.    3. The B1 field is generated by a resonant transmitter coil tuned exactly to the Larmor frequency and with its field direction perpendicular to B0.    4. The amplitude or duration of the B1 radio frequency pulse is adjusted typically to tip the nuclear magnetisation through 90 degrees from alignment with the direction of B0.    5. Following termination of the pulse, the nuclear magnetisation is in a plane orthogonal to the main magnetic field and is precessing at the Larmor frequency. The rotating nuclear magnetisation induces a weak signal in a resonant detection receiver coil (which can be the same coil as the transmitter coil) also tuned to the Larmor frequency. This signal from the nuclei is about a million times weaker than the signal originally transmitted by the radio frequency system.    6. The nuclear magnetisation undergoes a process of dephasing that results in a signal loss characterised by a decay time constant known as T2 (T2*).    7. Over a longer period the nuclear magnetization, free from the driving torque of the radio frequency pulse, re-aligns to its equilibrium position parallel to the direction of B0. This process is known as longitudinal magnetisation recovery and is characterised by a typical time constant T1.
In traditional 2D Fourier Transform MR imaging with phase encoding, following selection of a given slice, the next step is to encode the selected slice in order to obtain required in plane resolution and thereby to enable an image to be generated. To enable the slice to be encoded further gradients therefore have to by applied, these being known as the read (frequency) encode gradient and the phase encode gradient, these gradients being orthogonal to each other. Gradient switching, that is switching from read encode to phase encode and so on, of this kind is time consuming since it involves switching the read gradient on and then off and then switching the phase gradient on and then off etc. until the required image is acquired, processed and displayed accordingly.
In the known MR imaging approach described above there is therefore a problem in that if more than one frequency range is excited at a given time the detector 107 receives signals which the central processing unit 110 is unable to unscramble. In other words a superposition of images arises if an attempt is made to try and image more than one slice at a given time. However it is known to attempt to excite multiple slices simultaneously.
A first group of methods of acquiring slices in parallel, which have been widely used, is based around the idea of Hadamard encoding. Relevant disclosures in this respect include:                Souza S P, Szumowski J, Dumoulin C L, Plewes D P, Glover G. SIMA: Simultaneous multi-slice acquisition of MR images by Hadamard—encoded excitation, J Comput Assist Tomogr 1988:12:1026-30;        Muller S. Multifrequency selective RF pulses for multislice MR imaging, Magn Reson Med; 1988, 6:364:71; and        Glover G, Shimawaka A POMP (Phase offset multi-planar) imaging: a new high efficiency technique. Proc SMRM7th Annual Meeting, 1988, 241).        
These Hadamard methods use phase cycling over multiple excitations to present aliasing artefacts, thus lengthening overall scan time when signal to noise ratio is adequate to avoid averaging. Thus use of Hadamard pulses in this way usually requires a number of averages for phase cycling due to the need for image additions and subtractions to remove aliasing. These methods have been found to result in longer overall scan times and motion sensitivity. By motion sensitivity it is meant, in a medical imaging context for example, patient movement, since it is difficult in many circumstances to expect a patient to remain still for a sufficient time period whilst a given image is being acquired.
A further prior art method of obtaining multiple image slices is described by Larkman D J et al, J. Magn. Reson. Imaging 2001; 13(2): 313-317. Larkman's method concerns use of multicoil arrays for separation of signals from multiple slices excited simultaneously and although it permits simultaneous slice-imaging it utilizes a number of RF coils and a number of expensive receiver channels resulting in complex computer processing requirements.
A further prior art simultaneous slice method is disclosed by J B Weaver under the title “Simultaneous multislice acquisition of MR images”, in the journal Magn. Reson. Med. 1988; 8: 275-284. Weaver's method uses a slice selection gradient applied during readout to separate slices with frequency over sampling, but the method suffers from severe image distortion and blurring artefact due to oblique angulation of the image readout gradient.
A further parallel slice method concerns using a stepped magnetic field as is disclosed by Jakob P, Vath A and Haase A in German patent publication number DE 4216969. This method concerns a Simultaneous Excitation Simultaneous Acquisition Method known by the acronym “SESAM”. However the method disclosed does not incorporate any other form of parallel imaging.
Other well researched methods of reducing scan time are the parallel B1 coil array methods such as SMASH or SENSE (SENSitivity Encoding) developed recently by Sodickson et al. and Pruessman et al. based on earlier work. SENSE, for example, is a type of parallel imaging method concerning radio frequency sensitivity encoding with reduced phase views. Relevant references include:    1. Ra J B, Rim C Y. Fast imaging using sub-encoding using data sets from multiple detectors. Magn Reson Med. 1993; 30: 142-5.    2. Hutchison M, Raff U. Fast MRI data acquisition using multiple detectors. Magn Reson Med 1988; 6: 87-91.    3. Sodickson D K, Manning W J. Simultaneous acquisition of spatial harmonics (SMASH): fast imaging with radio frequency coil arrays. Magn Reson Med. 1997; 38(4): 591-603.    4. Pruessmann K P, Weiger M, Scheidegger M B, Boesiger P. SENSE: sensitivity encoding for fast MRI. Magn Reson Med. 1999 November; 42(5): 952-62.    5. Pruessmann K P, Weiger M, Bornert P, Boesiger P. Advances in sensitivity encoding with arbitrary k-space trajectories. Magn Reson Med. 2001; 46(4): 638-51.
These SMASH/SENSE type prior art methods use phase under sampling to reduce scan time in combination with reconstruction based on receiver coil sensitivity profiles to remove resulting fold-over phase encode aliasing artefacts. SENSE can also be used to produce parallel slice images, but with the need for many more receiver channels if in-plane encoding is also required (Larkman D J, Hajnal J V, Herlihy A H, Couts G A, Young I R and Ehnholm G. Use of multicoil arrays for separation of signal from multiple slices simultaneously excited. J Magn Reson Imaging. 2001; 13(2): 313-317).
In view of the above there is clearly a need to improve the known magnetic resonance apparatus and methods of MRI imaging so as to enable multiple slices through a given physical entity being imaged to be obtained simultaneously and thereby increase the speed of obtaining required images.
Similarly in some applications concerning multi-dimensional MR imaging there is a need to be able to read out images in a single imaging operation wherein time consuming switching of the one or more gradient fields is no longer required or at least is substantially reduced.